X-ray computer tomograph

ABSTRACT

An X-ray computer tomography apparatus is disclosed wherein a beam filter arrangement ( 26 ) that comprises regions ( 28, 30 ) of different filter material or/and different thickness profile of the filter material is arranged in the beam path preceding an examination subject. For multi-spectra beam hardening correction with estimation of the base material lengths, the computer tomography apparatus is fashioned for implementing projections of the examination subject given different effective material thickness or/and different effective material of the beam filter arrangement ( 26 ).

[0001] The invention is directed to an X-ray computer tomography apparatus with multi-spectra beam hardening correction.

[0002] The attenuation p that X-radiation generated by an X-ray source experiences in a transirradiated subject is measured in X-ray computer tomography. It is determined from the X-ray intensity I₀ incident onto the subject and from the intensity I that is registered in a detector arranged in the beam path following the subject, being determined according to the following equation:

p=−1n(I/I ₀)  (1).

[0003] In the case of mono-energetic radiation, the following applies for a homogeneous subject with the attenuation coefficient μ and the transirradiated subject thickness d:

p=μd  (2).

[0004] The X-ray attenuation thus increases linearly with the subject thickness.

[0005] In fact, however, an X-ray tube emits polychromatic X-radiation with the energy distribution S(E). The attenuation is then calculated according to the following equation:

  (3).

[0006] Even when the subject is homogeneous, the X-ray attenuation produced by the subject is thus no longer linearly dependent on the transirradiated subject thickness. Since μE usually decreases toward higher energies, the “energy center of gravity” shifts toward higher energies, namely all the more the greater the transirradiated subject thickness is. This effect is referred to as beam hardening.

[0007] In image reconstruction methods that are standard in CT technology, a linear change of the X-ray attenuation with the subject thickness is assumed for homogeneous subjects. The overall attenuation p of a beam on its path through a subject composed of partial subjects i with attenuation coefficient μ_(i) and thickness d_(i) then derives from:

  (4).

[0008] The deviations from this assumption caused by the beam hardening lead to data inconsistencies and, thus, to image errors. Typical image errors caused by beam hardening are key artifacts in large, homogeneous subjects and line or bar artifacts in CT images with a high proportion of bone or contrast agent. Current correction methods often have the principal goal or eliminating key artifacts and far-reaching bar artifacts in subjects with high attenuation, for instance in shoulder and pelvis exposures. These corrections usually ensue with what is referred to as polynomial correction, whereby a corrected attenuation value p_(c) is calculated from a detected measured attenuation value p_(M) by insertion into a polynomial with predetermined coefficients a_(n) according to the following equation:

  (5).

[0009] The coefficients a_(n) are acquired, for example, by measuring the attenuation values of uniform absorbers (for example, plexiglass bars) given N different thicknesses.

[0010] It has been shown that improved correction methods are needed for the correction of locally limited bar and line artifacts as well as unsharp bone-tissue transitions as particularly occur given skull exposures (another known bar artifact, for example, is what is referred to as the Hounsfield bar between the petrous bones). An approach has thereby proven beneficial wherein the length of the “base material” that the X-ray beam leading to a measured value has traversed in the body of the patient under examination is individually estimated for each measured attenuation value. In medical examinations, bone substance and soft tissue or, respectively, water, which has spectral attenuation properties similar to soft tissue, are usually selected as base materials. What is referred to as the two-spectra method, for example, is known from the pertinent literature for estimating the base material lengths traversed by an X-ray beam. Given this method, two measured values with respectively different spectral energy distribution of the X-ray, which is equivalent to a different average energy of the X-ray, are registered. Given known attenuation coefficients μ_(W)(E₁) and μ_(W)(E₂) of water at the average spectral energies E₁ and E₂ and μ_(K)(E₁) and μ_(K)(E₂) of bone at these average energies, the following, approximate estimate is possible for the measured attenuation values p(E₁) and p(E₂) obtained given there energies E₁ and E₂:

p(E ₁)=d _(W)·μ_(W)(E ₁)+d _(K)·μ_(K)(E ₁)  (6a)

p(E ₂)=d _(W)·μ_(W)(E ₂)+d _(K)·μ_(K)(E ₂)  (6b).

[0011] The water and bone lengths d_(W) and d_(K) can then be estimated from Eqq. (6) and (7) [sic].

[0012] A corrected measured value p_(c)(E₁) or, respectively, p_(c)(E₂) can now be respectively determined in the following way for the average spectral energies E₁ and E₂:

p _(c)(E ₁)=p(E ₁)+f _(E1)(d _(W) , d _(K))  (7a)

p _(c)(E ₂)=p(E ₂)+f _(E2)(d _(W) , d _(K))  (7b)

[0013] The correction values f_(E1) and f_(E2) are thereby taken from tables that were determined in advance either computationally or empirically for the average spectral energies E₁ and E₂.

[0014] Further information about the above two-spectra method can be found, for example, in the following publications:

[0015] 1) P. M. Joseph, R. D. Spittal, Journal of Computer Assisted Tomography, 1978, Vol.2, p.100;

[0016] 2) P. C. Johns, M. Yaffe, Medical Physics, 1982, Vol. 9, p.231;

[0017] 3) G. H. Glover, Medical Physics, 1982, Vol. 9, p.860;

[0018] 4) A. J. Coleman, M. Sinclair, Physics in Medicine and Biology, 1985, Vol. 30, No. 11, p.1251.

[0019] In order to register measured values at two different average energies in traditional CT apparatus, two successive revolutions of the X-radiator around the patient must be implemented. In the second revolution, work is performed with a different beam pre-filtering or with a different tube voltage than in the first revolution. What is disadvantageous about such a procedure, however, is that the measured results can exhibit inconsistencies due to patient movement or contrast agent flow.

[0020] Compared thereto, the invention provides an x-ray computer tomograph apparatus comprising

[0021] a radiator-detector arrangement that supplies projection measured values for each slice projection of an examination subject in allocation to a plurality of detection channels of the detector distributed over the entire projection region of this slice projection, each of said projection measured values being representative of the attenuation of the x-radiation in the respective detection channel produced by the examination subject, whereby the radiator-detector arrangement is fashioned for respectively supplying a plurality of at least two projection measured values given respectively different average energy of the x-radiation entering into the examination subject for multi-spectra beam hardening correction in allocation to each detection channel lying within the projection region of a slice projection;

[0022] an electronic evaluation and reconstruction unit connected to the radiator-detector arrangement that is fashioned for determining a beam hardening-corrected projection value for each of the projection measured values and of reconstructing a tomographic image of the examination subject upon employment of the corrected projection values; and

[0023] energy influencing means for influencing the average energy of the x-radiation entering into the examination subject.

[0024] It is inventively provided in this computer tomography apparatus that the energy influencing means comprise a beam filter arrangement arranged in the beam path preceding the examination subject that comprises different filter material or/and different thickness profile of the filter material for influencing the average energy of the x-radiation, and that the radio-detector arrangement is fashioned for respectively supplying a plurality of at least two projection measured values given respectively different filter material or/and different thickness profile of the filter material of the beam filter arrangement in allocation to each detection channel lying within the projection region of a slice projection.

[0025] The varying material or/and the varying material thickness of the beam filter arrangement make it possible to realize different average energies of the x-radiation entering into the examination subject with one and the same beam filter arrangement without having to change the beam filter arrangement. In particular, the projection measured values for the various average energies can be registered in immediate chronological proximity to one another, so that falsifying influences on the projection measured values due to contrast agent flow and physical movements on the part of the patient need not be feared. All projection measured values can then be registered in one revolution of the radiator of the radiator-detector arrangement.

[0026] Since a multi-spectra correction with estimate of the base material lengths will not be required in all examination scenarios, it is recommended that the beam filter arrangement be interchangeably mounted at the radiator-detector arrangement in order to also keep the employability of the computer tomography apparatus opened for other correction techniques as well.

[0027] The beam filter arrangement can be held in a simple way at a diaphragm carrier arranged radiator-proximate that carries a diaphragm arrangement for beam shaping of the x-radiation emitted by the radiator.

[0028] Given computer tomography apparatus having the possibility of a spring focus mode wherein the radiator of the radiator-detector arrangement is implemented with at least two spring foci between which it can be switched in alternation, the inventive solution can be utilized in such a way that the beam filter arrangement—in allocation to each of the spring foci—respectively comprises a region of different filter material or/and different thickness profile of the filter material, and such that the radiator-detector arrangement is fashioned for supplying a respective projection measured value for each of the spring foci in allocation to each detection channel lying within the projection region of a slice projection.

[0029] The evaluation and reconstruction unit can thereby be fashioned for determining an effective projection value by weighted summation from corrected projection values determined in allocation to respectively one of the detection channels and respectively allocated to one of the spring foci and for reconstructing the tomographic image upon employment of the effective projection values. In this way, the effect can be compensated that the projection measured values of a detection channel are registered at different spectra.

[0030] The spring focus mode, however, also leaves the possibility open of realizing slice projections of enhanced sampling density in that the evaluation and reconstruction unit is fashioned for reconstructing the tomographic image for a plurality of projection channels per slice projection that is equal to a multiple of the plurality of detection channels lying within the projection region of the respective slice projection corresponding to the plurality of spring foci, whereby the evaluation and reconstruction is fashioned for employing the corrected projection values determined in allocation to respectively one of the detection channels for all spring foci in the reconstruction of the tomographic image as corrected projection values of neighboring projection channels.

[0031] The inventive solution can also be advantageously utilized in computer tomography devices wherein the detector of the radiator-detector arrangement is implemented with a plurality of detector elements arranged in at least two lines lying above one another, an identical detection channel being allocated to their detector elements lying above one another in a respective column. In this case, the beam filter arrangement—in allocation to at least a sub-plurality of at least two detector elements of each column of detector elements lying within the projection region of a slice projection—can respectively comprise a region of different filter material or/and different thickness profile of the filter material, whereby the radiator-detector arrangement is then fashioned for supplying a respective projection measured value for each detector element from this sub-plurality of detector elements in allocation to each column of detector elements lying within the projection region of this slice projection.

[0032] In order to thereby avoid subjectively perceived changes between tomographic images based on the employment of different energy spectra that are reconstructed from the projection measured values of successive lines of detector elements, the evaluation and reconstruction unit can be fashioned for determining an effective projection value by weighted summation from the corrected projection values determined in allocation to respectively one of the columns and respectively allocated to one of the detector elements from the sub-plurality of detector elements and of reconstructing the tomographic image upon employment of the effective projection values.

[0033] The invention is explained in greater detail below on the basis of the attached drawings. Shown therein are:

[0034]FIG. 1 schematically, an inventive exemplary embodiment of CT scanner with spring focus mode;

[0035]FIG. 2 schematically, an inventive exemplary embodiment of a multi-line CT scanner; and

[0036]FIG. 3 schematically, a version of a radiation pre-filter for the CT scanners of FIGS. 1 and 2.

[0037] The CT scanner to be seen in FIG. 1 comprises an x-radiator 10 having two identical foci 12, 14 between which the x-radiator 10 can skip back and forth. Proceeding from each of the foci 12, 14, the x-radiator 10 can emit x-radiation onto the body 16 of a patient under examination fan-like in one plane. A detector arrangement 18 detects the radiation passing through the body 16. It comprises a plurality of detector elements 20 arranged next to one another on a circular arc in the direction of the fan angle, each of these detector elements 20 covering a part of the total projection region of the slice projection generated by irradiation of the body 16. Each of the detector elements 20 outputs an intensity measured signal that indicates the intensity of the incident radiation in the respective sub-region of the projection, outputting this to an electronic evaluation and reconstruction unit 22. The radiation intensity arriving in each individual sub-region of the projection is thus detected in its own detection channel. Using the incoming intensity measured signals, the evaluation and reconstruction unit 22 respectively determines an attenuation measured value that indicates the beam attenuation in the respective sub-region of the projection.

[0038] The radiator 10 is movable around the body 16 in a rotatory direction 24 and implements slice projections of the body 16 under a plurality of projection angles. Two slice projections are taken at each projection angle, one upon employment of the focus 12 and one upon employment of the focus 14.

[0039] In order to be able to implement a two-spectra beam hardening correction of the attenuation measured values obtained given the slice projections, the two slice projections taken at each projection angle are implemented with different average energies of the x-radiation entering into the body 16 under examination. To this end, a beam pre-filter 26 is arranged in the beam path of the x-radiation preceding the body 16, different average spectral energies for the two foci 12, 14 being capable of being set with said radiation pre-filter 26. The radiation pre-filter 24 [sic] has two filter regions 28, 30 that—in the illustrated exemplary embodiment—differ in view of their filter material given the same material thickness but, alternatively or additionally, can also have a different thickness profile. The radiation pre-filter 26 is arranged such that the filter region 28 is effective given employment of the focus 12, whereas the filter region 30 is effective given employment of the focus 14. The different filter material of the filter regions 28, 30 then produces the desired difference in the average spectral energies. Even though this cannot be seen in FIG. 1, the radiation pre-filter 26 is expediently curved such that the path that all individual rays of the ray fan beamed out by the radiator 10 traverse the radiation pre-filter 26 is approximately the same, so that an additional calibration of the detection channels can be avoided.

[0040] On the basis of the attenuation measured values acquired at the various average spectral energies, the evaluation and reconstruction unit 22 implements an estimate of the lengths of the base materials traversed by the x-rays in the body 16. Water and bone are considered below as base materials. It is thereby assumed that the base material lengths d_(W)(k) and d_(K)(k) of water and bone to be estimated for each individual detection channel k are approximately the same given the two projections that are implemented at each projection angle with the focus 12 on the one hand and with the focus 14 on the other hand. The following estimate then applies for the attenuation measured values p_(E1)(k) and p_(E2)(k) that the evaluation and reconstruction unit 22 determines for each detection channel k at the two average spectral energies E1 and E2:

p(k,E1)=d _(W)(k)μ_(W)(E1)+d _(K)(k)μ_(K)(E1)  (8a)

p(k,E1)=d _(W)(k)μ_(W)(E2)+d _(K)(k)μ_(K)(E1)  (8b)

[0041] The two unknowns d_(W)(k) and d_(K)(k) can be determined from this equation system. The evaluation and reconstruction unit 22 then calculates corrected attenuation values p_(c)(k, E1) and p_(c)(k, E2) given recourse to previously determined tables, from which it takes correction values f_(E1) and f_(E2) dependent on the values d_(W) and d_(K):

p _(c)(k,E1)=p(k,E1)+f _(E1)(d _(W)(k),d _(K)(k))  (9a)

p _(c)(k,E2)=p(k,E2)+f _(E2)(d _(W)(k),d _(K)(k))  (9b)

[0042] A tomograph image could then be reconstructed merely from the corrected attenuation values p_(c)(k, E1), as it also could be reconstructed only from the attenuation values p_(c)(k, E2). In both instances, a tomographic image would be reconstructed with a channel plurality N per slice projection that is equal to the plurality of detection channels formed by the detector arrangement 18, i.e. equal to the plurality of detector elements 20 lying next to one another in the direction of the fan angle in the respective overall projection region. One might be of the opinion that the same tomographic image derives in both instances. In fact, however, it is possible that image differences can be found that are based thereon that the measured values were registered given different average spectral energies. In order to compensate this effect, a tomographic image can be reconstructed from data that derive from a weighted summation of the attenuation values p_(c)(k, E1) and p_(c)(k, E2). In case of an averaging with identical weighting, one then obtains attenuation values p_(c)′(k) from the following:

p _(c)′(k)=0.5[p _(c)(k,E1)+p _(c)(k,E2)]  (10)

[0043] It is self-evident that a different weighting of the attenuation values p_(c)(k, E1) and p_(c)(k, E2) can also be undertaken whenever desired.

[0044] Alternatively, it is conceivable that a tomographic image be reconstructed with an increased channel plurality M in that the corrected attenuation values p_(c)(k, E1) and p_(c)(k, E2) are viewed as being the result of a single projection. Attenuation values p_(c)″(1) are thereby formed in the following way by interleaving the attenuation values p_(c)(k, E1) and p_(c)(k, E2):

p _(c)″(1=2k)=p _(c)(k,E1)  (11a)

p _(c)″(1=2k−1)=p _(c)(k,E2)  (11b)

[0045] whereby k−1, 2, . . . N. In this way, an attenuation value p_(c)″(1) is obtained for each channel 1 from a plurality M of projection channels that is twice as great as the plurality N of the detection channels formed by the detector arrangement 18.

[0046]FIGS. 2 and 3 shall now be referenced. Identical components or components having the same effect as in FIG. 1 are thereby provided with the same reference characters but supplemented by a lower case letter. In order to avoid repetitions, it is essentially only the differences compared to the exemplary embodiment of FIG. 1 that shall be explained. For the rest, the above description of FIG. 1 is referenced.

[0047] The CT scanner shown in FIG. 2 is what is referred to as a multi-line scanner that comprises detector elements 20 a in a plurality of lines lying above one another in the direction of a z-axis 32 a. The z-axis 32 a thereby corresponds to the feed axis along which the patient 16 a is moved through CT scanner. In the illustrated exemplary embodiment of FIG. 2, the detector arrangement 18 a comprises four such lines of detector elements 20 a; the lines are referenced Z1, Z2, Z3 and Z4. The x-radiation 10 a is fashioned for implementing a slice projection of the body 16 a at each projection angle for each detector line. In the present exemplary case, thus, four slice projections following one another in the direction of the Z-axis can be registered at each projection angle. All detector elements 20 a lying above one another in a column supply measured signals at these four slice projections that—given a view in the direction of the fan angle of the radiation band emitted by the radiator 10 a from a focus 34 a at each slice projection—can be respectively allocated to the same detection channel. In order to implement a multi-spectra attenuation measurement given the multi-line scanner of FIG. 2, the radiation pre-filter 26 a attached focus-proximate comprises regions of different thickness profile or/and different filter material in the direction of the z-axis. Dependent on the design of the radiation pre-filter 26 a, a different average spectral energy of the x-radiation can be obtained for each detector line or respectively the same average spectral energy can be obtained for groups of detector lines. In the exemplary case of FIG. 2, the radiation-pre-filter 26 a has a thickness that varies in z-direction. It is thereby symmetrically designed in z-direction such that the respectively same filter effect is achieved for the inner detector lines Z2 and Z3.

[0048] Because the radiation pre-filter 26 a extends over all detector lines in z-direction and no discontinuous transitions of the filter effect occur, the artifacts susceptibility of the tomographic images reconstructed for the individual detector lines is decidedly slight even given imprecise adjustment or mechanical movement of the pre-filter or given unavoidable gravitational or thermal z-movement of the focus 34 a during the revolution of the radiator 10 a around the patient 16 a. A steady curve of the filter effect of the radiation pre-filter 26 a is therefore preferred in z-direction.

[0049] The radiation pre-filter 26 a can be mounted in a diaphragm box 36 a (indicated with broken lines) wherein a diaphragm arrangement 38 a is accommodated that serves the purpose of beam shaping of the x-radiation emitted by the radiator 10 a in z-direction and in the direction of the fan angle. Together with a traditional, auxiliary pre-filter arrangement (not shown in greater detail), the radiation pre-filter 26 a can thereby releasably mounted on a common changing device, so that it is removable as needed and can only be brought into use for specific purposes (for example, exposures of the base of the skull).

[0050] The following description of a multi-spectra beam hardening correction for the CT scanner shown in FIG. 2 having four detector lines Z1 through Z4 can be transferred without further ado onto other CT scanner having a different number of lines. A two-spectra correction shall be considered first for the case of the radiation pre-filter 26 a according to FIG. 2 that is symmetrical relative to the middle of the detector. For the line pair Z1 and Z2, the different effective thickness of the radiation pre-filter 26 a supplies spectra having different average quantum energy E₁ or, respectively, E₂. The same is true for the line pair Z3 and Z4. Given the assumption of approximately identical estimated values for d_(W) and d_(K) for both detector lines of the respective line pair, a corrected attenuation value can be respectively calculated. The following then applies for the channels k of a projection p_(i)(k, Ej), registered at the effective energy Ej (j=1,2) in the detector line i (i=1,2,3,4):

p ₁(k,E1)=d _(W1)(k)μ_(W)(E1)+d _(K1)(k)μ_(K)(E1)  (12a)

p ₂(k,E2)=d _(W1)(k)μ_(W)(E2)+d _(K1)(k)μ_(K)(E2)  (12b)

p ₃(k,E2)=d _(W2)(k)μ_(W)(E2)+d _(K2)(k)μ_(K)(E2)  (12c)

p ₄(k,E1)=d _(W2)(k)μ_(W)(E1)+d _(K2)(k)8355 2 _(K)(E1)  (12d)

[0051] Using the Equations ( 12 a) through (12d), the base material lengths d_(W1) and d_(W2) can now be determined for water (soft tissue) as well as d_(K1) and d_(K2) for bone. Employing correction factors f_(Ej) (d_(Wv)(k), d_(Kυ)(k)) (v=1,2) that, for example, are taken from pre-calculated tables, corrected attenuation values p_(ci)(,Ej) for all detector lines i (i=1,2,3,4) can then be determined in the following way:

p _(c1)(k,E1)=p₁(k,E1)+f _(E1)(d _(W1)(k),d _(K1)(k))  (13a)

p _(c2)(k,E2)=p₂(k,E2)+f _(E2)(d _(W1)(k),d _(K1)(k))  (13b)

p _(c3)(k,E2)=p₃(k,E2)+f _(E2)(d _(W2)(k),d _(K2)(k))  (13c)

p _(c4)(k,E1)=p₄(k,E1)+f _(E1)(d _(W2)(k),d _(K2)(k))  (13d)

[0052] When a respective tomographic image for each of the different detector lines is reconstructed from the corrected attenuation values p_(ci)(k,Ej), then it can occur that the image impression subjectively changes between the lines of the line pair Z1 and Z2 and between the lines of line pair Z3 and Z4 due to the respectively different average spectral energy. In order to avoid this effect, the separately corrected attenuation values p_(ci)(k, Ej) of the different detector lines can be averaged to form attenuation values P_(cq)(k, E_(eff)) of two effective detector lines q (q=1,2):

p _(ci)(k,E _(eff))=0.5[p _(c1)(k,E1)+p _(c2)(k,E2)]  ( 14 a)

p _(c2)(k,E _(eff))=0.5[p _(c3)(k,E2)+p _(c4)(k,E1)]  ( 14 b)

[0053] A common tomographic image is reconstructed for the detector lines Z1 and Z2 from the effective attenuation values p_(c1)(k, E_(eff)), whereas a common tomographic image for the detector lines Z3 and Z4 is reconstructed from the effective attenuation values p_(c2)(k, E_(eff)). The same effective energy E_(eff) is thereby to be allocated to the effective attenuation values p_(c1)(k, E_(eff)) and P_(c2)(k, E_(eff)).

[0054] The projections can also be combined with different weighting of the individual attenuation values p_(ci)(k, Ej) in order to intensify the image effectiveness of one of the two energies E₁ and E₂:

p _(c1)(k,E _(eff))=g ₁ p _(c1)(k,E1)+g ₂ p _(c2)(k,E2)  (15a)

p _(c2)(k,E _(eff))=g ₂ p _(c3)(k,E2)+g ₁ p _(c4)(k,E1)  (15b)

[0055] whereby g₁+g₂=1 applies.

[0056] Over and above this, a utilization of the four detector lines Z1 through Z4 at four different energy spectra is conceivable, even though the outlay then rises in view of calculating time and table generation. In such a four-spectra correction, further base materials could be taken into consideration, for example iodine-containing water solutions. Beam hardening errors that occur given exposures of the brain and other body parts as a consequence of the employment of iodine-containing contrast agent could then be eliminated. FIG. 3 shows a possible structure of a radiation pre-filter 26 b in order to realize a four-spectra mode of the four-line scanner of FIG. 2. The radiation pre-filter 26 b fabricated, for example, of titanium exhibits a thickness that changes uniformly in z-direction over all detector lines.

[0057] Given four detector lines under consideration, the following equation system derives for determining the base material lengths d_(W) of water, d_(K) of bone and d_(X) and d_(Y) of two further materials X and Y. It must thereby be assumed that the four base material lengths d_(W), d_(K), d_(X), d_(Y) to be determined are at least approximately constant for all detector lines under consideration.

p ₁(k,E1)=d _(W)(k)μ_(W)(E1)+d _(K)(k)μ_(K)(E1)++d _(X)(k)μ_(X)(E1)+d _(Y)(k)μ_(Y)(E1)  (16a)

p ₂(k,E2)=d _(W)(k)μ_(W)(E2)+d _(K)(k)μ_(K)(E2)++d _(X)(k)μ_(X)(E2)+d _(Y)(k)μ_(Y)(E2)  (16b)

p ₃(k,E3)=d _(W)(k)μ_(W)(E3)+d _(K)(k)μ_(K)(E3)++d _(X)(k)μ_(X)(E3)+d _(Y)(k)μ_(Y)(E3)  (16 c)

p ₄(k,E4)=d _(W)(k)μ_(W)(E4)+d _(K)(k)μ_(K)(E4)++d _(X)(k)μ_(X)(E4)+d _(Y)(k)μ_(Y)(E4)  (16d)

[0058] When the four base material lengths have been determined on the basis of this equation system, a correction factor f_(Eυ)(d_(W)(k), d_(K)(k), d_(X)(k), d_(Y)(k)) can be taken from previously determined tables for all energies E_(υ)(υ=1,2,3,4). The corrected attenuation values P_(cυ)(k, E_(υ)) are then calculated analogous to the two-spectra method:

p _(cυ)(k,E _(υ))=p _(υ)(k,E _(υ))+f _(Eυ)(d _(W)(k),d _(K)(k),d _(X)(k),d _(Y)(k))  (17)

[0059] whereby v=1, 2, 3, 4. In this case, too, an image reconstruction can individually cause a subjectively buried image impression from line to line for each of the detector lines. As a result of weighted combination of all four corrected projections, this can again be avoided:

p _(c)(k,E _(eff))=Σ[_(υ=1,2,3,4]) [g _(υ) p _(cυ)(k,Eυ)]  (18)

[0060] whereby Σ_([υ=1,2,3,4]) gυ=1 applies. A common tomographic image from these effective attenuation values is then reconstructed for all four detector lines.

[0061] It must be added that the invention, of course, can also be employed given multi-line scanners having spring focus mode, whereby it is then possible to employ a radiation pre-filter that comprises thickness variations or/and material variations both in the direction of the fan angle as well as in z-direction. 

1. X-ray computer tomograph apparatus comprising a radiator-detector arrangement (10, 18) that supplies projection measured values for each slice projection of an examination subject (16) in allocation to a plurality of detection channels of the detector (18) distributed over the entire projection region of this slice projection, each of said projection measured values being representative of the attenuation of the x-radiation in the respective detection channel produced by the examination subject (16), whereby the radiator-detector (10, 18) arrangement is fashioned for respectively supplying a plurality of at least two projection measured values given respectively different average energy of the x-radiation entering into the examination subject for multi-spectra beam hardening correction in allocation to each detection channel lying within the projection region of a slice projection; an electronic evaluation and reconstruction unit (22) connected to the radiator-detector arrangement (10, 18) that is fashioned for determining a beam hardening-corrected projection value for each of the projection measured values and of reconstructing a tomographic image of the examination subject (16) upon employment of the corrected projection values; and energy influencing means (26) for influencing the average energy of the x-radiation entering into the examination subject (16), characterized in that the energy influencing means (26) comprise a beam filter arrangement (26) arranged in the beam path preceding the examination subject that comprises regions (28, 30) of different filter material or/and different thickness profile of the filter material for influencing the average energy of the x-radiation, and in that the radio-detector (10, 18) arrangement is fashioned for respectively supplying a plurality of at least two projection measured values given respectively different filter material or/and different thickness profile of the filter material of the beam filter arrangement (26) in allocation to each detection channel lying within the projection region of a slice projection.
 2. Computer tomography arrangement according to claim 1, characterized in that the beam filter arrangement (26) is interchangeably mounted at the radiator-detector arrangement (10, 18).
 3. Computer tomography arrangement according to claim 1 or 2, characterized in that beam filter arrangement (26 a) is held at a diaphragm carrier (36 a) arranged radiator-proximate that carries a diaphragm arrangement (38 a) for beam shaping of the x-radiation emitted by the radiator (10).
 4. Computer tomography arrangement according to one of the claims 1 through 3, characterized in that the radiator (10) of the radiator-detector arrangement (10, 18) is implemented with at least two spring foci (12, 14) between which it can be switched in alternation; in that the beam filter arrangement (26)—in allocation to each of the spring foci (12, 14)—respectively comprises a region (28, 30) of different filter material or/and different thickness profile of the filter material, and in that the radiator-detector arrangement (10, 18) is fashioned for supplying a respective projection measured value for each of the spring foci (12, 14) in allocation to each detection channel lying within the projection region of a slice projection.
 5. Computer tomography arrangement according to claim 4, characterized in that the evaluation and reconstruction unit (22) is fashioned for determining an effective projection value by weighted summation from corrected projection values determined in allocation to respectively one of the detection channels and respectively allocated to one of the spring foci (12, 14) and for reconstructing the tomographic image upon employment of the effective projection values.
 6. Computer tomography arrangement according to claim 4 or 5, characterized in that the evaluation and reconstruction unit (22) is fashioned for reconstructing the tomographic image for a plurality of projection channels per slice projection that is equal to a multiple of the plurality of detection channels lying within the projection region of the respective slice projection corresponding to the plurality of spring foci (12, 14), whereby the evaluation and reconstruction (22) is fashioned for employing the corrected projection values determined in allocation to respectively one of the detection channels for all spring foci (12, 14) in the reconstruction of the tomographic image as corrected projection values of neighboring projection channels.
 7. Computer tomography arrangement according to one of the claims 1 through 6, characterized in that the detector (18 a) of the radiator-detector arrangement (10 a, 18 a) is implemented with a plurality of detector elements (20 a) arranged in at least two lines (Z1 through Z4) lying above one another, an identical detection channel being allocated to their detector elements (20 a) lying above one another in a respective column; in that the beam filter arrangement (26 a)—in allocation to at least a sub-plurality of at least two detector elements (20 a) of each column of detector elements (20 a) lying within the projection region of a slice projection—respectively comprises a region of different filter material or/and different thickness profile of the filter material; and in that the radiator-detector arrangement (10 a, 18 a) is fashioned for supplying a respective projection measured value for each detector element (20 a) from this sub-plurality of detector elements (20 a) in allocation to each column of detector elements (20 a) lying within the projection region of this slice projection.
 8. Computer tomography arrangement according to claim 7, characterized in that the evaluation and reconstruction unit (22 a) is for determining an effective projection value by weighted summation from the corrected projection values determined in allocation to respectively one of the columns and respectively allocated to one of the detector elements (20 a) from the sub-plurality of detector elements (20 a) and of reconstructing the tomographic image upon employment of the effective projection values. 